Object imaging using diffuse light

ABSTRACT

Imaging tumors using diffuse light. An imaging system includes a source of diffuse light for generating oscillatory diffuse photon density waves to illuminate an object, a detector for detecting diffuse photon density waves interacting with the object, and a computer interfaced with the detector for processing data corresponding to the photon density waves detected to determine at least a position of the object. In one embodiment, the turbid medium and the object have associated therewith at least one diffusion coefficient and the diffuse photon density waves which illuminate the object refract around the object as a result of their interaction with it, thereby producing a distorted wavefront that allows the computer to construct an image of the object. In another embodiment, a fluorescent object produces re-radiated diffuse photon density waves which allow the object to be imaged.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a United States National Phase Application ofPCT/US94/12486 filed Oct. 31, 1994 which is a continuation of U.S.application Ser. No. 08/145,466 filed Oct. 29, 1993, (now abandoned).

FIELD OF THE INVENTION

This invention relates generally to imaging of objects. Morespecifically, this invention relates to methods and apparatus forimaging objects using diffuse light.

BACKGROUND OF THE INVENTION

Techniques for imaging objects have been used for nearly a century inthe medical arts for diagnosing and understanding the myriad diseasesand maladies that afflict the human body. Imaging techniques have alsofound use in such diverse fields as radio astronomy, sonar, radar andother fields which require information about an object which is notreadily visible to the naked eye and therefore not easily examined.Medical imaging techniques include, for example, X-ray imaging, positronemission tomography (PET), ultrasound imaging and the well knownmagnetic resonance imaging (MRI).

In all of the imaging techniques mentioned above, narrow band frequencyradiation illuminates the object of interest to produce reflected oremitted radiation which is then gathered from the object by a detector.The reflected or emitted radiation is then processed by an imagingalgorithm to obtain useful information about the object.

In medical applications, the use of ionizing radiation in imaging, forexample with X-rays, involves significant health risks to a patient whenthe patient is exposed to the radiation for prolonged periods of time orin multiple imaging schemes. Furthermore, certain of these imagingtechniques undesirably involve the use of invasive procedures which areboth costly and painful. Yet other techniques such as MRI do not yieldconsistently useful clinical results.

There has thus arisen in the medical imaging art an interest indeveloping non-invasive, safe and relatively fast techniques which cantake advantage of the natural scattering of visible and infrared lightthrough media containing objects to be imaged. Techniques using diffuselight could be used in conjunction with other imaging schemes such asX-ray imaging or MRI to produce highly useful clinical images fordiagnostic purposes.

Much of the progress in imaging with diffusive light has focused onballistic techniques using lasers. With these techniques, an intensepulsed laser illuminates a sample. By time gating photons that have beenscattered only a few times and rejecting all other photons, the opticalabsorption of the medium and objects found therein can be mapped. Thistechnique works best when the allowed time window is short and thephotons deviate the least from their "ballistic" trajectory.Unfortunately, the transmittal intensity of unscattered photonsdiminishes exponentially with increasing sample thickness.

Because of the limitations of ballistic imaging, it is difficult toobtain high quality images of relatively thick objects with low powerlasers. Examples of ballistic imaging techniques are disclosed in K. M.Yoo, F. Lie and R. R. Alfano, Optics Letters, Vol. 16, p. 1068 (1991),and in D. A. Benaron and D. K. Stevenson, Science, Vol. 259, p. 1463(1993).

A second technique for imaging using diffuse light is optical phasemodulation. Phase modulation techniques have permitted the location ofsingle absorbers using low power, continuous wavelength lasers. Inaccordance with these techniques, an amplitude modulated source createsphoton density waves that acquire anomalous phase shifts due to theabsorber. For the case of a single absorber, the distortions are readilyinterpreted; however for a more complicated object a general analysis isrequired.

An example of imaging with diffuse light is disclosed in U.S. Pat. No.5,119,815, Chance where scattered light was applied to a biologicalimaging application. The Chance patent discloses a technique for solvingthe diffusion equation for a homogeneous medium to obtain the overalloptical absorption characteristics. This was possible for thehomogeneous medium because the long time limit of the logarithmicderivative of the detected intensity yields the absorptioncharacteristics directly. Thus the absorption characteristics foruniform structures may be obtained with the methods and apparatusdisclosed in the Chance patent.

Still other attempts to image with diffuse light are disclosed in U.S.Pat. No. 5,070,455, Singer et al. In the Singer et al. system, lightintensities are measured at many sensor positions (pixels), initialvalues of absorption or scattering coefficients are assigned at eachpixel, and then a new set of intensities at each pixel is calculated.The calculated intensities are compared to the real intensities, and theintensity differences are used to generate a subsequent interaction ofabsorption or scattering values for each pixel.

The methods described in Singer et al. usually require many iterationssince the absorption or scattering values may not converge rapidly.Furthermore, the Singer et al. system utilizes cumbersome Monte-Carlostatistical techniques which consume large amounts of processing timewithout guaranteeing computational success. Singer et al.'s methods mayalso produce false local minima providing misleading results for theabsorption characteristics.

Thus prior imaging techniques using diffuse light for scattering fail tosolve a long-felt need in the art for robust imaging techniques whichcan produce reliable images in biological systems. Solution of theaforementioned problems has heretofore eluded the medical imaging art.

SUMMARY OF THE INVENTION

The aforementioned deficiencies in the imaging art are overcome bymethods and apparatus provided in accordance with the present inventionwhich provide imaging of objects in turbid media using diffuse light.

In a preferred embodiment of the invention an imaging system comprisessource means for generating oscillatory diffuse photon density waves toilluminate the object, detection means for detecting diffuse photondensity waves produced as a result of the diffuse photon density wavesinteracting with the object, and processing means interfaced with thedetection means for processing data corresponding to the photon densitywaves detected by the detection means to determine at least a positionof the object. The turbid medium and the object have associatedtherewith at least one diffusion coefficient and the diffuse photondensity waves which illuminate the object diffract around or refractthrough the object as a result of their interaction with it, therebyproducing a distorted wavefront such that after the detection meansdetects the distorted wavefront the processing means determines thediffusion coefficient of the turbid medium and the object. Morepreferably, the processing means constructs phase and amplitude contourscorresponding to propagation of the distorted wavefront and furtherdetermines at least the position of the object from the phase andamplitude contours, thereby imaging the object.

In yet a further preferred embodiment of imaging systems provided inaccordance with the invention, display means are interfaced with theprocessing means for displaying the image of the object, the sourcemeans comprises at least one laser, and the detection means comprises anoptical fiber interfaced with a photomultiplier tube.

Further aspects of the invention provide imaging of a fluorescent objectsuch that the diffuse photon density waves having a first wavelengthcause the object to fluoresce, thereby producing re-radiated diffusephoton density waves having a second wavelength such that after thedetection means detects the re-radiated diffuse photon density waves,the processing means can image the object.

In the embodiment of the invention where fluorescent, re-radiateddiffuse photon density waves are detected, the source means preferablycomprises a plurality of lasers oriented around the object whichalternately irradiate the object with the diffuse photon density wavesof the first wavelength to cause the object to fluoresce. The detectionmeans comprises an optical fiber that is placed in proximity to theobject and a photomultiplier tube interfaced to the optical fiber. Theimaging system further comprises switch means interfaced with each ofthe plurality of lasers for alternately and sequentially turning on andoff each laser, and radio frequency driving means interfaced through theswitch means with the lasers for driving the lasers to produce thediffuse photon density waves of the first wavelength.

Alternatively, the imaging system comprises a plurality of lasers eachhaving a spatial location with respect to the object. Each laser is morepreferably modulated at all times during imaging at a differentfrequency in a frequency range around a specified frequency, therebyproducing a power spectrum associated with each spatial location aroundthe object. Analysis means are provided interfaced with the detectionmeans and the processing means for analyzing the power spectrumsassociated with each spatial location to determine the position of theobject.

In still a further preferred embodiment of imaging systems which takeadvantage of fluorescent, re-radiated diffuse photon density waves, thesource means comprises a phased-array. The phased-array, preferablycomprises at least two lasers that are substantially one hundred andeighty degrees out of phase with each other, thereby producing thediffuse photon density waves having the first wavelength which interferedestructively to produce an amplitude null line and a substantially onehundred and eighty degree phase shift across the null line equidistantfrom the lasers. By scanning the null line with the phased-array theprocessing means can produce an image of the object in the turbidmedium.

Systems and methods provided in accordance with the present inventionwill provide efficient imaging of tumors and other maladies that effecthuman tissue. These systems will also prove to be much more economicalto build when compared to prior imaging systems, since they will notrequire the complex machinery that is associated with prior imagingsystems such as MRI. Furthermore, since the methods and apparatusprovided in accordance with the present invention utilize diffuse lightfor imaging, more reliable images of tumors will be presented to themedical diagnostician or clinician so that cancerous tumors and otherinhomogeneities in the tissue will be detected earlier and more readily.This holds the promise of saving lives and reducing the overall costs ofmedical care.

The features, objects and advantages of the invention will be understoodby those with skill in the art from the following detailed descriptionof the preferred embodiments thereof when read in conjunction with thedrawings which are first described briefly below.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a schematic representation of an imaging system using diffuselight wherein scattering experiments with diffuse photon density wavesoccurs.

FIG. 2 depicts measured constant contours of diffuse photon densitywaves propagating through a turbid medium.

FIG. 3 depicts measured constant phase contours of diffuse photondensity waves propagating through turbid media and refracting across aplane boundary between two turbid media with different light diffusioncoefficients.

FIG. 4 is a graph of phase contours of diffuse photon density wavespropagating through turbid media after refracting across a cylindricalboundary separating two turbid media wherein lensing effects areobserved.

FIG. 5 is a schematic representation of a "time-sharing" system forimaging diffuse photon density waves re-radiated from a fluorescentobject.

FIG. 6A is a graph of phase contours of the disturbance produced by afiber point source in a turbid medium.

FIG. 6B is a graph of the absorption and emission characteristics of afluorescent dye contained in a spherical shell within the turbid medium.

FIG. 6C is a graph of measured constant amplitude phase contours fordiffuse photon density waves emitted from a source into a turbid medium,and re-radiated diffuse photon density waves.

FIG. 7 is a schematic representation of a "frequency-encoded" system forimaging diffuse photon density waves re-radiated from a fluorescentobject.

FIGS. 8A and 8B are schematic representations of "phased-array" scanningsystems for imaging re-radiated diffuse photon density waves and graphsof the object's position in the turbid medium and a null line associatedwith the phased-array.

DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS

Propagating disturbances are produced in a dense, turbid mediumcontaining objects when amplitude modulated light sources are introducedinto the turbid medium. In biological and medical imaging applications,human tissue, such as breast tissue containing tumors, is such a turbidmedium containing objects to be imaged. In accordance with the presentinvention, an oscillatory light source introduces diffuse photon densitywaves (hereinafter referred to as "DPDW") in the turbid medium. Whilethe results described herein have been achieved with respect to severalexperimental apparatus to be described hereafter, those with skill inthe art will immediately recognize that the techniques and apparatusdisclosed are readily applicable to imaging human body tissue which isinfected with tumors or other maladies that will appear asinhomogeneities in the tissue which is a turbid medium.

The inventors of the subject matter herein claimed and disclosed havediscovered that DPDW can be used in at least two ways to performimaging. In a first preferred embodiment, DPDW are introduced into aturbid medium and refracted through and diffracted around objects in themedium, thereby producing a distorted wavefront which can then beanalyzed to yield useful information concerning at least a position andsize of the object around which the DPDW have been refracted.

In a further preferred embodiment, the inventors have determined that byintroducing fluorescence to the object, photons in the re-radiated DPDWwill be emitted from the object at a shifted wavelength with respect tothe original DPDW and the re-radiated DPDW then analyzed to determine atleast the position of the fluorescent objects in the turbid medium. Thisis particularly useful for imaging of tumors since in prior, non-diffuselight imaging techniques fluorescent dyes which respond to various formsof radiation have typically been introduced into tumors, therebyyielding information about the location and size and nature of the tumorunder examination.

Thus, imaging tumors and objects in turbid media with diffuse light willin the future produce significant and effective images for diagnosticand clinical purposes.

I. The Basic Theory of DPDW Propagation in a Turbid Medium

The basic mechanics of wave propagation in a turbid medium have beenexplored by the inventors when performing imaging in accordance with thepresent invention. DPDW are scalar, over-damped, traveling waves oflight energy density, denoted U(r,t). They propagate through any mediumin which the transport of light energy density, U, is governed by the"diffusion equation," which is ##EQU1## where D is the "diffusioncoefficient" for the medium. This diffusion equation holds for anon-absorbing medium. Some examples of optically turbid media includedense suspensions of micrometer-sized spheres, human tissue, paints,foams, and Intralipid (a mixture of water, soybean oil and egg yolk).The introduction of amplitude modulated light into a turbid mediumproduces a macroscopic ripple of brightness that is microscopicallycomposed of individual photons undergoing random walks. The disturbancesarise whenever the diffusive system is driven by an oscillating source.

The oscillatory part of the solution for an infinite, non-absorbent,homogeneous turbid medium in the presence of a point source located atthe origin follows the form:

    U(r,t)=(A/Dr)(exp {-kr})(exp {i(kr-ωτ})

where A is a constant, r is the radial distance from the origin, D isthe diffusion coefficient for light in the medium, ω is the sourcemodulation frequency, and k=(ω/2D)^(1/2). Although the wave is veryrapidly attenuated, it has a well-defined wavelength, amplitude andphase at all points. Interestingly, the wavelength can be altered bymodifying D or ω.

When absorption is present, a similar solution to Equation (1) can beobtained, but the real and imaginary parts of the wave vector k aredifferent and depend explicitly on the sample absorption length (as wellas the photon random walk step for inverse scattering factor). Themacroscopic disturbance obeys a Helmholtz equation, and therefore hasmany properties that are normally associated with conventionalelectromagnetic radiation. Thus, DPDW display refraction, diffraction,and interference properties similar to those observed with conventionalelectromagnetic radiation propagating through media. By examining theenergy contours produced as a result of the interaction of the DPDW withthe objects in the medium according to the techniques of the presentinvention, the objects can be imaged to obtain useful clinical data.

II. Imaging Using Refracted DPDW

In a preferred embodiment, an apparatus for imaging objects in a turbidmedium is shown schematically in FIG. 1. The object 10 may containsmaller objects such as tumors which must be characterized for clinicalpurposes. For experimental purposes, the dense turbid medium imaged isIntralipid, a polydisperse suspension of particles having an averagediameter of ˜0.4 μm, but a relatively wide range of sizes (i.e., from˜0.1 to ˜1.1 μm). By changing the solution concentration of theIntralipid, it is possible to vary the light diffusion coefficient, D,of the medium. The photon transport mean free path l* was about 0.2 cmin a 0.5% concentrated solution.

To perform imaging of the smaller objects in the Intralipid, a largefish tank 20 (30 cm×30 cm×60 cm) is filled with this material. When theabsorption is very small, the suspensions in the Intralipid are dilute,and therefore the diffusion coefficient is inversely proportional to theIntralipid concentration.

In a preferred embodiment, a source fiber 30 and detector fiber 40 (both˜4 mm in diameter) are placed in proximity to object 10. The sourcelight 30 is preferably derived from a 3-mW diode laser 35 operating atabout 816 nm. The diode laser 35 is amplitude modulated at 200 MHz bydriver 50, and the position of the source fiber 30 is fixed. Thedetector fiber 40 is in detecting proximity to the object 10 and isfurther connected to a photomultiplier tube 60 on its other end. Inorder to facilitate phase and amplitude measurements, both the referencefrom the driver 50 and the detected signal are down-converted to 25 kHzby heterodyning with a second oscillator 70 to 200.025 MHz. Mixers 80preferably combine the frequencies to produce the 25 kHz signals shown.The low-frequency signals are then measured using a lock-in amplifier90.

In a more preferred embodiment, the phase shift (and ac amplitude) ofthe detected light is measured with respect to the source 30 at eachpoint on a 0.5-cm square planar grid throughout the sample. Constantphase contours are then determined by linear interpolation of the griddata. The sensitivity of the apparatus shown in FIG. 1 is about 10⁵.Since the signal amplitude decays by <e⁻² π in one wavelength, the rangeof the apparatus is limited to slightly more than one wavelength.Nevertheless, it is possible to clearly distinguish the essentialphysical phenomena of the DPDW scattering through the object 10 withthis apparatus.

The results obtained by imaging with the system of FIG. 1 demonstratethat a "diffusional index of refraction" exists for DPDW, and that it ispossible to manipulate the diffusional index of refraction bycontrolling the photon diffusion coefficients (D) of adjacent turbidmedia. This is of considerable importance in biological systems, wherethe natural curvature of organs such as the brain, heart, or kidney,together with changes of scattering and absorption as in the grey-whitematter transition of the brain, can lead to significant modifications ofthe trajectories of diffuse photons.

Results for a ˜0.5% concentrated homogeneous turbid medium areillustrated in FIG. 2. Constant phase contours are shown at 20°intervals about the source. The contours are approximately circular, andtheir radii can therefore be extrapolated back to the source 30. In theinsert of FIG. 2, the phase shift and the quantity ln|rU_(ac) (rt)| as afunction of radial distance from the source are plotted. From thesemeasurements it is possible to determine that the wavelength of the DPDWis 11.2 cm, the photon transport mean free path is ˜0.2 cm, and thephoton absorption length is ˜52 cm in ˜0.5% Intralipid at 22° C. In thiscase, the photon absorption can be attributed almost entirely to water.

Referring to FIGS. 3 and 4, refraction of the DPDW is shown in threeways: (1) when a plane, acrylic partition or (2) a cylindrical acrylicpartition is placed in the tank to simulate a boundary, and (3) when nopartition is in place so that the medium is homogeneous. In FIG. 3,constant phase contours exist at about every 20°. When the planeboundary 45 is introduced, the lower medium 55 has a concentration,c_(l), ˜1.0% and light diffusion coefficient D_(l). The upper medium 65has a concentration, c_(u), ≈0.25% and light diffusion coefficientD_(u). The contours below boundary 45, shown at the 4 cm position on they-axis, are just the homogeneous media contours without refraction.These contours are obtained before the partition is introduced. Thecontours above the boundary 45 are derived from the DPDW that aretransmitted into the less concentrated medium.

Theory predicts that the wavelength in the less dense medium, λ_(u), is14.8 cm, and should be greater than the wavelength of the DPDW in theincident medium, λ_(l), which is 8.17 cm. The ratio of the twowavelengths should equal the ratio of the diffusional indices ofrefraction of the two media. As is predicted, λ_(u) =λ_(l) (D_(l)/D_(u))^(-1/2) ˜λ_(l) (c_(l) /c_(u))^(1/2).

Theory also predicts that the apparent source position S_(i), as viewedfrom within the upper medium 65, should be shifted from the real sourceposition, S_(o) =4.0±0.2 cm, by a factor λ_(l) /λ_(u) =0.55. This isverifiable from the data of FIG. 3 since, using the radii from the fullcontour plots, the apparent source position is shifted from 4.0±0.2 to2.0±0.25 cm.

In accordance with the invention, FIG. 3 explicitly demonstrates Snell'slaw for DPDW. This can be seen by following the ray from S_(o) to thepoint A at the boundary 45, and then into the upper medium 65. The rayin the lower medium 55 makes an angle θ_(i) =14° with respect to thesurface normal. The upper ray is constructed in the standard way betweenthe apparent source position S_(i), through the point A on boundary 45,and into the medium 65 above the boundary. The upper ray isperpendicular to the circular wave fronts in the less dense medium 55,and makes an angle θ_(t) =26.6° with respect to the boundary normal.Therefore, it can be seen graphically that sinθ_(i) /sinθ_(t)=0.54≈λ_(l) /λ_(u), so that Snell's law accurately describes thepropagation of DPDW across boundary 45.

The inventors have also determined that by using a circular boundary(shown generally at 75 in FIG. 4) to separate the two turbid media, thecurvature of the DPDW can be altered in analogy with a simple lens inoptics. Referring to FIG. 4, two semi-infinite media are separated bycurved boundary 75, and the medium 55 on the right is more concentrated.The constant phase contours of the transmitted wave exhibit a shorterwavelength, and are clearly converging toward some image point to theright of the boundary. The medium 65 on the left (λ_(l)) has anIntralipid concentration of ˜0.1%, and the medium 55 on the right(λ_(r)) has a concentration of ˜1.6%.

The wavelength ratio is measured to be λ_(r) /λ_(l) =3.8±0.3. The curvedsurface 75 has a radius R=9.0±0.4 cm. The position of source 30 is S_(o)=9.4±0.3 cm. The image position is determined to be S_(i) =12±2 cm. Thisresult deviates somewhat from the well-known paraxial result fromgeometrical optics for imaging by a spherical refracting surface. Thedeviation is primarily a result of spherical aberration. However, inaccordance with the invention, the curvature of the wave fronts isreversed after traversing the circular boundary 75.

The results shown in FIGS. 2, 3 and 4 which are obtained from theapparatus of FIG. 1 show that it is possible to exert substantialcontrol over the transport of diffuse light in dense random media. Theinventors have clearly demonstrated that the index of refraction of DPDWin such a medium depends on the photon diffusion coefficient or randomwalk step of the photons in the medium. This allows imaging ofinhomogeneities in the medium by examining the DPDW, a result that hasnot heretofore been achieved in the art.

It should be noted that when deriving Equation (1), the differentialform of Fick's law and photon flux conservation principles were used. Ina preferred embodiment of imaging by examining the scattering anddiffraction refraction of DPDW, the effects of absorption of the DPDWcan be ignored, and it can be assumed that the time it takes for lightto travel a single random walk step is much shorter than the modulationperiod. In this case, the oscillatory part of the light energy densityU.sub.ω (r) obeys the Helmholtz equation, i.e., (V² +k²)U.sub.ω (r)=0.The only significant difference of DPDW propagation in comparison toconventional wave phenomena is that k² =i(ω/D), and therefore k iscomplex.

The spatial part of U_(ac) (r,t) in Equation (1) is simply the Green'sfunction solution of the Helmholtz equation with the appropriate k.Therefore, some of the basic theorems that apply to solutions of theHelmholtz equation will apply to DPDW propagation in a turbid medium.For example, a Kirchoff integral can be constructed for these wavesusing the Green's function solution. This provides a formal method bywhich to calculate the wave amplitude and phase at various distancesfrom a "diffracting" aperture as has been discussed in this Part II. Tothe extent that the Kirchoff integral embodies the basic Huygens-Fresnelprinciple, contributions of different elements of a scattering surfacearising from damped, spherical point sources will be observed. This alsoimplies that the focusing of DPDW will have the same limitations due todiffraction as in the case of light propagating in a standard opticalmedium. Thus, imaging of tumor-like inhomogeneities in tissue can beaccomplished in accordance with the present invention by examining therefractive, diffractive and scattering properties of DPDW incident tothe tissue.

III. Imaging by Examining Re-Radiated DPDW

The inventors of the subject matter herein claimed and disclosed havealso discovered that examining re-radiated DPDW from a fluorescentinhomogeneity in a turbid medium provides methods of imaginginhomogeneities in the medium. Many types of tumors which occur in thehuman body comprise inhomogeneities which can be made to fluoresce afterbeing irradiated with DPDW, and therefore, the inventors have determinedthat re-radiated DPDW provide an excellent means of locating tumors.

In order to observe re-radiated DPDW and image an object, the modelbiological material Intralipid is used. An amplitude modulated 200 MHzor 50 MHz laser diode ˜3 mW, 780 nm is fiber-coupled into the medium,and another optical fiber is used to detect diffuse photons as afunction of position within the medium. Using the standard heterodynetechniques discussed earlier, the phase and amplitude of the DPDW in themedium can be observed.

A preferred apparatus for imaging inhomogeneities in turbid media bydetecting re-radiated DPDW is shown schematically in FIG. 5. Theabsorber/radiator shown at 100 preferably comprises a shell filled withIntralipid at a concentration of 0.4 mg/L Indocyanine green dye which isless than one tenth of the concentration commonly used in human subjectsto test hepatic function. In this embodiment, multiple sources 110 and asingle detector 120 are used to determine the center of the fluorescentobject 100.

Computer processor 130 preferably outputs a control signal to switch 140which controls the sequential activation of each of sources 110. Adetector 150 which most preferably comprises a photomultiplier tube isinterfaced both to detector fiber 120 and computer processor 130. In afurther preferred embodiment, detector 150 is a Hamamatsu R928 or R1645uphotomultiplier tube which also comprises a high voltage power supplythat ensures adequate gain to computer processor 130 so the computerprocessor 130 can process the data received from detector fiber 120. Adisplay device 160, preferably a CRT screen, is interfaced to computerprocessor 130 and outputs an image 170 of object 100.

To obtain image 170, a fluorescent object 100 is irradiated withmultiple sources 110 of DPDW. It is then preferred to measure theamplitude (or phase) of the re-radiated light. The partial amplituderesulting solely from source i, is dependent on the i^(th) source-objectseparation, the quantum efficiency of the dye in the object 100, and theobject 100-detector 120 separation. Therefore: ##EQU2## where ξ is thequantum efficiency of the dye, r_(i) is the position of the i^(th)source, r_(o) is the position of the object center, r_(d) is thedetector position, and k (k') is the wave vector magnitude of the DPDWat 780 nm (830 nm).

The individual sources 110 are separately turned on and off, and there-radiated amplitude for each object separation is measured. Since thesource positions r_(i), and the detector position r_(d) are known, it ispossible to estimate the object's position by finding the value of r_(o)that gives the best agreement with the measured ratio |U_(i) |/|U_(j) |.Generally, three sources 110 are necessary to localize the object in twodimensions. However, it is more preferable to use four or more sourcesto improve the signal-to-noise ratio of the imaging system. Using foursources 110 as shown in FIG. 5, it is possible to localize the center ofthe 1 cm spherical object 100 to within 0.4 cm. This two-dimensionallocalization is easily extended to three dimensions, as will berecognized by those with skill in the art.

Referring to FIG. 6A, constant phase contours of the disturbanceproduced by the fiber point source located at the origin of the systemare illustrated. In a more preferred embodiment, re-radiated DPDW areobtained by filling a spherical glass shell with the absorbing dyeIndocyanine green, and then illuminating the sphere with DPDW in theIntralipid solution. The dye is preferably chosen to absorb radiation atthe source wavelength of 780 nm, and very soon thereafter re-radiatephotons at a red-shifted energy, 830 nm. Because the dye has a lifetimeof less than 1 nsec compared to the 5 nsec period of the source, there-radiated energy is in the form of a DPDW at the red-shifted energy.

The absorption and emission characteristics of the dye are shown in FIG.6B wherein the inset is the chemical formula of the dye material. TheIntralipid solution surrounding the obstacle has a concentration of 0.1%giving a source diffuse photon density wavelength of ˜18 cm. A pointsource at the origin generates the incident DPDW. Constant amplitudecontours of the incident wave in the presence of the obstacle are shownin FIG. 6C (dashed lines). In a more preferred embodiment, two spectralfilters centered on 830 nm (Schott glass filters, RG830) are provided toenable the incident and re-radiated DPDW to be separated.

In FIG. 6C, the measured constant amplitude contours of the wave at 830nm are shown as solid lines, and the measured incident amplitudecontours at 780 nm are shown as dashed lines. The dashed linesdemonstrate the DPDW character of the re-radiated waves. The re-radiatedwave originates from within the absorbing object. As can be seen fromthe contours, the DPDW wavelength at 830 nm is somewhat longer than theDPDW wavelength at 780 nm. This is a function of the relatively largerdiffusion coefficient for 830 nm light in Intralipid. Thus, a type offluorescence of DPDW has occurred and the inhomogeneity is convertedinto a source of secondary DPDW.

An alternative embodiment to the "time-sharing apparatus" for imagingusing re-radiated DPDW described by FIG. 5 is a "frequency encodedapparatus" shown in FIG. 7. In this embodiment, each source 120 ismodulated by a slightly different modulation frequency, f_(i), (i.e.,f_(i) =200.00+k (0.01) MHz) and is kept on at all times. In this way, amodulation power spectrum is associated with each spatial location inthe sample. By measuring the power spectrum of the re-radiated light asa function of modulation frequency with a spectrum analyzer 180, theposition of object 100 can be determined using essentially the sameanalysis as described above with respect to the apparatus of FIG. 5.This type of imaging is similar to MRI in that a frequency spectrumemitted by an object is converted to the spatial location of the object.

A third, alternative embodiment for imaging fluorescent objects thatre-radiate DPDW is shown in FIGS. 8A and 8B. The apparatus of FIGS. 8Aand 8B is qualitatively different than the time-sharing apparatus andfrequency encoded apparatus of FIGS. 5 and 7 respectively, since it usesa scanning phased-array 190 and a single detector 200. Preferably, thephased-array 190 comprises two sources 85 and 95 that are substantially180° out of phase with each other and that emit DPDW which interferedestructively to produce an amplitude null and a sharp 180° phase shiftacross a curve that describes this family of points called the "nullline," shown schematically at 210.

By placing a detector 200 on the null line 210, and then moving anabsorbing object 220 from one side of the null line 210 to the other,the object 220 will preferentially absorb light from the nearest source,and therefore distort the null line. When the absorber 220 is also are-radiator, the complimentary effect is seen, that is, the objectre-radiates more light derived from the closest source. In bothmeasurements, the phase of the detected DPDW will undergo a 180° shiftas the absorber crosses the original, undisturbed null line 210. Theseeffects are demonstrated by the graphs of FIG. 8A. The re-radiated lightdisplays a sharper phase transition, and a deeper amplitude null thanthat of the incident light, and the re-radiated wave exhibits acomplimentary change as discussed above.

Interestingly, in the embodiment of FIG. 8A, the phase shift of there-radiated light is always the same, independent of detector position.Thus, it is possible to hold the absorber 220 and the detector 200stationary, and scan the null line 210 which will produce essentiallythe same results. Scanning the null line 210 can be achieved bymechanically translating the two sources 85 and 95 together.

Referring to FIG. 8B this effect is demonstrated for re-radiated light.A sharp phase shift at the location of the re-radiator 220 and null line210 is detected. The phase transition of the incident light occurs nearthe position of detector 200. With knowledge of the position of the nullline 210 as a function of time, a one-dimensional localization of there-radiator can be achieved. By performing three scans down threeperpendicular axes, three-dimensional localization of re-radiator 220can be achieved. It is envisioned that the phased-array apparatus ofFIGS. 8A and 8B will be simple to build and implement. Furthermore,imaging apparatus described herein will generally be much moreeconomical to build than other imaging systems currently in use, such asMRI systems, PET systems, and CAT scanners.

Therefore, it has been demonstrated that imaging of inhomogeneities inturbid media using DPDW is both possible and practical with the methodsand apparatus disclosed and claimed herein. By studying the scatteringand distortion of the DPDW by inhomogeneities found therein, theinhomogeneities can be located and their sizes determined for clinicalpurposes. Alternatively, a fluorescent inhomogeneity can be examined byanalyzing the re-radiated DPDW emitted therefrom. In accordance with theinvention, it is also possible to combine the dual techniques ofexamining the refractive, diffractive, scattering, and fluorescentproperties of the inhomogeneities to image the objects.

The techniques and apparatus of the present invention provideparticularly powerful clinical tools for imaging, locating and sizingtumors in human tissue. Furthermore, the methods and apparatus claimedand disclosed herein should prove to be both economical and efficientimaging tools for clinical diagnosis.

There have thus been described certain preferred embodiments of methodsand apparatus for imaging objects using diffuse light provided inaccordance with the present invention. While preferred embodiments havebeen described and disclosed, it will be recognized by those with skillin the art that modifications are within the true spirit and scope ofthe invention. The appended claims are intended to cover all suchmodifications.

What is claimed is:
 1. A system for imaging an object in a turbid mediumusing diffuse light comprising:source means for generating oscillatorydiffuse photon density waves to illuminate the object; detection meansfor detecting diffuse photon density waves produced as a result of thediffuse photon density waves interacting with the object; processingmeans interfaced with the detection means for processing datacorresponding to the photon density waves detected by the detectionmeans to determine at least a position of the object; said turbid mediumand the object have associated therewith at least one diffusioncoefficient and the diffuse photon density waves which illuminate theobject diffracts around or refracts through the object as a result oftheir interaction with it, thereby producing a distorted wavefront suchthat after the detection means detects the distorted wavefront theprocessing means determines the diffusion coefficient of the turbidmedium and the object.
 2. The system recited in claim 1 wherein theprocessing means constructs phase and amplitude contours correspondingto propagation of the distorted wavefront and further determines atleast the position of the object from the phase and amplitude contours,thereby imaging the object.
 3. The system recited in claim 2 furthercomprising display means interfaced with the processing means fordisplaying the position of the object.
 4. The system recited in claim 3wherein the source means comprises at least one laser and the detectionmeans comprises an optical fiber interfaced with a photomultiplier tube.5. The system recited in claim 1 wherein the object is fluorescent andthe diffuse photon density waves which have a first wavelength cause theobject to fluoresce, thereby producing re-radiated diffuse photondensity waves having a second wavelength such that after the detectionmeans detects the re-radiated diffuse photon density waves, theprocessing means can image the object.
 6. The system recited in claim 5wherein the source means comprises a plurality of lasers oriented aroundthe object which alternately irradiate the object with the diffusephoton density waves of the first wavelength to cause the object tofluoresce.
 7. The system recited in claim 6 wherein the detection meanscomprises an optical fiber that is placed in proximity to the object anda photomultiplier tube interfaced to the optical fiber.
 8. The systemrecited in claim 7 further comprising switch means interfaced with eachof the plurality of lasers for alternately and sequentially turning onand off each laser and radio frequency driving means interfaced throughthe switch means with the lasers for driving the lasers to produce thediffuse photon density waves of the first wavelength.
 9. The systemrecited in claim 8 further comprising display means interfaced with theprocessing means for displaying the position of the object produced bythe processing means.
 10. The system recited in claim 5 wherein thesource means comprises a plurality of lasers each having a spatiallocation with respect to the object and each laser being modulated atall times during imaging at a different frequency in a frequency rangearound a specified frequency, thereby producing a power spectrumassociated with each spatial location around the object.
 11. The systemrecited in claim 10 further comprising analysis means interfaced withthe detection means and the processing means for analyzing the powerspectrums associated with each spatial location to determine theposition of the object.
 12. The system recited in claim 11 furthercomprising display means interfaced with the processing means fordisplaying the image of the object.
 13. The system recited in claim 5wherein the source means comprises a phased-array.
 14. The systemrecited in claim 13 wherein the phased-array comprises at least twolasers that are substantially one hundred and eighty degrees out ofphase with each other, thereby producing the diffuse photon densitywaves having the first wavelength which interfere destructively toproduce an amplitude null line and a one hundred and eighty degree phaseshift across the null line.
 15. The system recited in claim 14 whereinthe phased-array scans the null line so that the processing means canproduce the image.
 16. A method of imaging an object in a turbid mediumusing diffuse light comprising the steps of:illuminating the object withoscillatory diffuse photon density waves; detecting diffuse densitywaves which are produced as a result of the diffuse photon density wavesinteracting with the object; and determining at least the position ofthe object in the turbid medium by analyzing the detected diffuse photondensity waves, wherein the turbid medium and the object have associatedtherewith at least one diffusion coefficient and the diffuse photondensity waves which illuminate the object refract around the object as aresult of the interaction with it, thereby producing a distortedwavefront which is analyzed to determine the diffusion coefficient ofthe turbid medium and the object.
 17. The method recited in claim 16wherein the determining step comprises the steps of constructing phasecontours corresponding to propagation of the distorted wavefront anddetermining at least the position of the object from the phase contours,thereby imaging the object.
 18. The method recited in claim 17 furthercomprising the step of displaying the image of the object.
 19. Themethod recited in claim 16 wherein the object is fluorescent and thediffuse photon density waves which have a first wavelength cause theobject to fluoresce, thereby producing re-radiated diffuse photondensity waves having a second wavelength such that detecting there-radiated diffuse photon density waves the object to be imaged. 20.The method recited in claim 19 wherein the illuminating step comprisesthe step of alternately irradiating the object with at least two sourcesof diffuse photon density waves of the first wavelength to cause theobject to fluoresce.
 21. The method recited in claim 19 wherein theilluminating step comprises the step of continuously illuminating theobject with at least two sources of diffuse photon density waves eachhaving a spatial location with respect to the object, wherein eachsource has associated therewith a frequency in a frequency range arounda specified frequency, thereby producing a power spectrum associatedwith each spatial location around the object.
 22. The method recited inclaim 19 wherein the illuminating step comprises the steps ofirradiating the object with a phased-array and scanning an amplitudenull line to produce the image of the object.
 23. A system for imagingan object in a turbid medium comprising:source means for illuminatingthe object with oscillatory diffuse photon density waves of a firstspecified wavelength, whereby the object will fluoresce re-radiatediffuse photon density waves of a second wavelength after beingilluminated with the oscillatory diffuse photon density waves of thefirst specified wavelength; detection means for detecting the re-radiatediffuse photon density waves of the second wavelength; and processingmeans interfaced with the detection means for processing datacorresponding to the re-radiated diffuse photon density waves of thesecond wavelength to determine at least the position of the object inthe turbid medium.
 24. The system recited in claim 23 wherein the sourcemeans comprises a plurality of lasers oriented around the object whichalternately irradiated the object with the diffuse photon density wavesof the first wavelength to cause the object to fluoresce.
 25. The systemrecited in claim 24 wherein the detection means comprises an opticalfiber that is placed in proximity to the object and a photomultipliertube interfaced to the optical fiber.
 26. The system recited in claim 25further comprising switch means interfaced with each of the plurality oflasers for alternately and sequentially turning on and off each laserand radio frequency driving means interfaced through the switch meanswith the lasers for driving the lasers to produce the diffuse photondensity waves of the first wavelength.
 27. The system recited in claim26 further comprising display means interfaced with the processing meansfor displaying the image of the object produced by the processing means.28. The system recited in claim 23 wherein the source means comprises aplurality of lasers each having a spatial location with respect to theobject and each laser is modulated at all times during imaging at adifferent frequency in a frequency range around a specified frequency,thereby producing a power spectrum associated with each spatial locationaround the object.
 29. The system recited in claim 28 further comprisinganalysis means interfaced with the detection means and the processingmeans for analyzing the power spectrums associated with each spatiallocation to determine the position of the object.
 30. The system recitedin claim 29 further comprising display means interfaced with theprocessing means for displaying the image of the object.
 31. The systemrecited in claim 23 wherein the source means comprises a phased-array.32. The system recited in claim 31 wherein the phased-array comprises atleast two lasers that are substantially one hundred and eighty degreesout of phase with each other, thereby producing the diffuse photondensity waves having the first wavelength which interfere destructivelyto produce an amplitude null line and a one hundred and eighty degreephase shift at points equidistant from the lasers.
 33. The systemrecited in claim 32 wherein the phased-array scans the null line so thatthe processing means can produce the image.
 34. A method of imaging anobject in a turbid medium comprising the steps of:illuminating theobject with oscillatory diffuse photon density waves of a firstspecified wavelength; allowing the object to fluoresce, therebyreradiating a diffuse photon density waves having a second wavelength;detecting the re-radiated diffuse photon density waves having the secondwavelength; and analyzing the photon density waves having the secondwavelength to determine at least the position of the object based on itsimage in the turbid medium.
 35. The method recited in claim 34 whereinthe illuminating step comprises the step of alternately irradiating theobject with at least two sources of diffuse photon density waves of thefirst wavelength to cause the object to fluoresce.
 36. The methodrecited in claim 34 wherein the illuminating step comprises the step ofcontinuously illuminating the object with at least two sources ofdiffuse photon density waves each having a spatial location with respectto the object, wherein each source has associated therewith a frequencyin a frequency range around a specified frequency, thereby producing apower spectrum associated with each spatial location around the object.37. The method recited in claim 34 wherein the illuminating stepcomprises the steps of irradiating the object with a phased-array andscanning an amplitude null line to produce the image of the object. 38.A system for imaging an object in a turbid medium wherein the turbidmedium and object have a diffusion coefficient associated therewithcomprising:source means for illuminating the object with oscillatorydiffuse photon density waves, whereby a distorted wavefront of diffusephoton density waves is created by refraction of the oscillatory diffusephoton density waves refracting around the object; detection means fordetecting the distorted wavefront of diffuse photon density wavesrefracted around the object; processing means interfaced with thedetection means for constructing phase contours corresponding topropagation of the distorted wavefront of diffuse photon density wavesto determine at least the position of the object in the turbid medium;and display means interfaced with the processing means for displayingthe image of the object.
 39. The system recited in claim 38 wherein thesource means comprises at least one laser and the detection meanscomprises an optical fiber interfaced with a photomultiplier tube.
 40. Amethod of imaging an object in a turbid medium wherein the turbid mediumand the object have a diffusion coefficient associated therewithcomprising the steps of:illuminating the object with oscillatory diffusephoton density waves; allowing the diffuse photon density waves torefract around the object throughout the turbid medium thereby producinga distorted wavefront of diffuse photon density waves; detecting thedistorted wavefront of diffuse photon density waves; analyzing thedistorted wavefront of diffuse photon density waves to determine atleast the position of the object in the turbid medium utilizing thesteps of constructing phase contours corresponding to propagation of thedistorted wavefront and determining at least the position of the objectfrom the phase contours, thereby imaging the object; and displaying theimage of the object.